The present embodiments relate to calibrating a counting digital X-ray detector.
X-ray systems are used for imaging for diagnostic examination purposes and for interventional procedures (e.g., in cardiology, radiology and surgery). X-ray systems 16, as shown in FIG. 1, have an X-ray tube 18 and an X-ray detector 17 (e.g., jointly mounted on a C-arm 19), a high-voltage generator for generating the tube voltage, an imaging system 21 (e.g., including at least one monitor 22), a system control unit 20 and a patient table 23. Biplane systems (e.g., having two C-arms) are likewise employed in interventional radiology. Generally, flat-panel X-ray detectors find application as X-ray detectors in many fields of medical X-ray diagnostics and intervention (e.g., in radiography, interventional radiology, cardioangiography), but also in therapeutic treatment applications for imaging within the context of monitoring and irradiation planning or mammography.
Flat-panel X-ray detectors in use today may be integrating detectors and are based mainly on scintillators having light that is converted into electrical charge in photodiode arrays. The electrical charge is read out (e.g., row by row) via active control elements. FIG. 2 schematically shows the layout of a currently used indirectly converting flat-panel X-ray detector, having a scintillator 10, an active readout matrix 11 made of amorphous silicon with a plurality of pixel elements 12 (e.g., with photodiode 13 and switching element 14) and drive and readout electronics 15 (see, e.g., M. Spahn, “Flat detectors and their clinical applications,” Eur Radiol. (2005), 15: 1934-1947). Depending on beam quality, the quantum efficiency for a CsI-based scintillator having a layer thickness of, for example, 600 mm lies between about 50% and 80% (see, e.g., M. Spahn, “Flat detectors and their clinical applications,” Eur Radiol. (2005), 15: 1934-1947). By this, the spatial frequency dependent DQE(f) (“detective quantum efficiency”) is upwardly limited, and for typical pixel sizes of, for example, 150 μm to 200 mm and for the spatial frequencies of 1 lp/mm to 2 lp/mm, of interest to the applications lies significantly below that. In order to enable new applications (e.g., dual-energy, material separation), but also to increase quantum efficiency further, the potential of counting detectors or of energy-discriminating counting detectors mainly based on direct-converting materials (e.g., CdTe or CdZTe=CZT) and contacted application-specific integrated circuits (ASICs) (e.g., implemented in CMOS technology) is being investigated to an increasing extent.
An exemplary design of such counting detectors is shown in FIG. 3. X-ray radiation is converted in the direct converter 24 (e.g., CdTe or CZT), and the charge carrier pairs generated are separated by way of an electrical field generated by a common top electrode 26 and a pixel electrode 25. The charge generates a charge pulse in one of the pixel-shaped pixel electrodes 26 of the ASIC 27. The charge pulse corresponds in height to the energy of the X-ray quantum and, if lying above a defined threshold value, is registered as a count event. The threshold value serves to differentiate an actual event from electronic noise or, for example, also to suppress k-fluorescence photons in order to avoid multiple counts. The ASIC 27, a corresponding section of the direct converter 24 and a coupling between direct converter 24 and ASIC 27 (e.g., using bump bonds 36 in the case of direct-converting detectors) in each case form the detector module 35 having a plurality of pixel elements 12. The ASIC 27 is arranged on a substrate 37 and connected to peripheral electronics 38. A detector module may also have one or more ASICs and one or more subsections of a direct converter, chosen as requirements dictate in each case.
The general layout of a counting pixel element 12 is shown schematically in FIG. 5. The electrical charge is collected via the charge input 28 in the pixel element and amplified there with the aid of a charge amplifier 29 and a feedback capacitor 40. In addition, the pulse shape may be adjusted in a shaper (e.g., filter) at the output (not shown). An event is counted such that a digital memory unit 33 (e.g., counter) is incremented by one if the output signal is above a selectable threshold value. This is verified via a discriminator 31. In principle, the threshold value may also be permanently predefined by an analog device, though generally the threshold value is applied via a digital-to-analog converter (DAC) 32 and is thus variably adjustable within a certain range. The threshold value may either be set pixel by pixel locally, as shown, via the discriminator 31 (e.g., local discriminator) and ASIC 32 (e.g., local ASIC) or else globally for a plurality of/all pixel elements via, for example, a global discriminator and ASIC. The counter status of the digital memory unit 33 may be read out via a drive and readout unit 38. FIG. 6 shows a corresponding schematic layout for an entire array of counting pixel elements 12 (e.g., 100×100 pixel elements of 180 mm each and a drive and readout unit 38). In this example, the array would have a size of 1.8×1.8 cm2. For large-area detectors (e.g., 20×30 cm2), as shown in FIG. 4, for example, a plurality of detector modules 35 are connected together (e.g., in this example, 11×17 would produce approximately the desired surface area) and detector modules 35 are connected via the common peripheral electronics, such as, for example, a drive and readout unit 38. Through silicon via (TSV) technology, indicated by the reference numeral 37 in FIG. 4, is used, for example, for realizing the connection between ASIC 27 and peripheral electronics in order to provide the modules are arranged side by side as tightly as possible in a four-sided array.
In the case of counting and energy-discriminating X-ray detectors, two, three (e.g., as shown in FIG. 7, with the reference numeral according to FIG. 5) or more threshold values are introduced, and the level of the charge pulse, corresponding to the predefined threshold values (e.g., discriminator threshold values), is classified into one or more of the digital memory units (e.g., counters). The X-ray quanta counted in a particular energy range may be obtained by forming the difference between the counter contents of two corresponding counters. The discriminators may be set, for example, with the aid of digital-to-analog converters either for the entire detector module or pixel by pixel within given limits or ranges. The counter contents of the pixel elements are read out module by module in succession via a corresponding readout unit. This readout process requires a certain amount of time during which it is not possible to continue counting without error.
Basically, various architectures are possible for implementing digital-to-analog converters (DAC). Two examples are described below.
1. Each pixel has a digital-to-analog converter with sufficiently fine analog graduation per bit and sufficiently high bit depth, so that both expected variations may be corrected, and the threshold value desired in each case (e.g., X-ray energy range) may be covered.
2. There is a “global” digital-to-analog converter per X-ray detector or per detector module (or just a few), which generates a common voltage (or current) for all pixel elements and thus makes provision for a global “rough threshold value”. In each pixel, there is also an additional local digital-to-analog converter that generates a further pixel-internal common voltage (current) that in combination with the global voltage (current) generates a pixel-specific total voltage (total current).
In a design of the type, the increments of the energies for global digital-to-analog converters and local digital-to-analog converters may be configured either as equal in width (e.g., 1 keV/bit) or with different widths (e.g., rather more roughly for the global digital-to-analog converter and finer for the local digital-to-analog converters). In this case, the local digital-to-analog converters may have a sufficiently high bit depth in order to cover the rougher increments of the global digital-to-analog converter. As an example of the case described, the global digital-to-analog converter has an increment of 2 keV/bit (or corresponding voltage differences or current differences) with a 6-bit depth (e.g., values between 0 and 126 keV may be covered). Correspondingly, the local digital-to-analog converters have 0.5 keV/bit and, for example, a 5-bit depth configured to compensate for the local fluctuations of the global digital-to-analog converter at the pixel element in a range of 16 keV. The example presupposes a linear behavior of the global digital-to-analog converter and the local digital-to-analog converters, as otherwise a correspondingly more generously dimensioned design will be necessary. A favorable choice of bit depths and energy increments is dependent on pixel and module design, X-ray detector material properties (e.g., CdTe), the clinical applications, and other factors.
Counting detectors with adjustable discriminator threshold values exhibit problems such as the following. The discriminator threshold values may vary from pixel element to pixel element. However, since only X-ray quanta having X-ray energy that lies above the threshold values are counted, different portions of the energy spectrum are counted, resulting in “threshold value noise”. This type of noise is a particular characteristic of counting and energy-discriminating counting detectors. The discriminator threshold values are modified via DAC values. These digital values are initially not assigned to certain physical energies. A calibration enabling an assignment of threshold values and energies in keV is therefore to be provided. The design of the X-ray detector may have different sizes of pixel elements (e.g., smaller effective pixel elements at the edges of detector modules compared with pixel elements that are arranged centrally on the detector module), such that differently high count rates are to be expected for geometric reasons. The detector material may also have detector material defects (e.g., Te inclusions, structure limits, field profiles or other). Such effects may lead to structural or “fixed pattern”-like noise. The radiation field, too, may not be homogeneous over the entire surface area of the X-ray detector, since an X-ray source may be punctiform, and therefore, the X-ray flux may vary depending on the distance separating tube focus and detector pixel. The heel effect may also lead to locally different radiation profiles due to direction-dependent reabsorption of the generated X-ray radiation in the tube anode on the X-ray detector.